Non-invasive photoelectric pulse oximetry has been previously described in U.S. Pat. Nos. 4,407,290, 4,266,554, 4,086,915, 3,998,550, 3,704,706, European Patent Application No. 102,816 published Mar. 13, 1984, European Patent Application No. 104,772 published Apr. 4, 1984, European Patent Application No. 104,771 published Apr. 4, 1984, and PCT International Publication WO86/05674 published October 9, 1986. Pulse oximeters are commercially available from Nellcor Incorporated, Hayward, Calif., U.S.A., and are known as, for example, Pulse Oximeter Model N-100 (herein "N-100 oximeter") and Model N-200 (herein "N-200 oximeter").
Pulse oximeters typically measure and display various blood flow characteristics including but not limited to blood oxygen saturation of hemoglobin in arterial blood, volume of individual blood pulsations supplying the flesh, and the rate of blood pulsations corresponding to each heartbeat of the patient. The oximeters pass light through human or animal body tissue where blood perfuses the tissue such as a finger, an ear, the nasal septum or the scalp, and photoelectrically sense the absorption of light in the tissue. The amount of light adsorbed is then used to calculate the amount of blood constituent being measured.
The light passed through the tissue is selected to be of one or more wavelengths that is absorbed by the blood in an amount representative of the amount of the blood constituent present in the blood. The amount of transmitted light passed through the tissue will vary in accordance with the changing amount of blood constituent in the tissue and the related light absorption.
For example, the N-100 oximeter is a microprocessor controlled device that measures oxygen saturation of hemoglobin using light from two light emitting diodes ("LED's"), one having a discrete frequency of about 660 nanometers in the red light range and the other having a discrete frequency of about 925 nanometers in the infrared range. The N-100 oximeter microprocessor uses a four-state clock to provide a bipolar drive current for the two LED's so that a positive current pulse drives the infrared LED and a negative current pulse drives the red LED to illuminate alternately the two LED's so that the incident light will pass through, e.g., a fingertip, and the detected o transmitted light will be detected by a single photodetector. The clock uses a high strobing rate, e.g., one thousand five hundred cycles per second, to be easily distinguished from other light sources. The photodetector current changes in response to the red and infrared light transmitted in sequence and is converted to a voltage signal, amplified, and separated by a two-channel synchronous detector--one channel for processing the red light waveform and the other channel for processing the infrared light waveform. The separated signals are filtered to remove the strobing frequency, electrical noise, and ambient noise and then digitized by an analog to digital converter ("ADC"). As used herein, incident light and transmitted light refers to light generated by the LED or other light source, as distinguished from ambient or environmental light.
The light source intensity may be adjusted to accommodate variations among patients' skin color, flesh thickness, hair, blood, and other variants. The light transmitted is thus modulated by the absorption of light in the variants, particularly the arterial blood pulse or pulsatile component, and is referred to as the plethysmograph waveform, or the optical signal. The digital representation of the optical signal is referred to as the digital optical signal. The portion of the digital optical signal that refers to the pulsatile component is labeled the optical pulse.
The detected digital optical signal is processed by the microprocessor of the N-100 oximeter to analyze and identify optical pulses corresponding to arterial pulses and to develop a history as to pulse periodicity, pulse shape, and determined oxygen saturation. The N-100 oximeter microprocessor decides whether or not to accept a detected pulse as corresponding to an arterial pulse by comparing the detected pulse against the pulse history. To be accepted, a detected pulse must meet certain predetermined criteria, for example, the expected size of the pulse, when the pulse is expected to occur, and the expected ratio of the red light to infrared light of the detected optical pulse in accordance with a desired degree of confidence. Identified individual optical pulses accepted for processing are used to compute the oxygen saturation from the ratio of maximum and minimum pulse levels as seen by the red wavelength compared to the maximum and minimum pulse levels as seen by the infrared wavelength, in accordance with the following equation: ##EQU1## wherein BO1 is the extinction coefficient for oxygenated hemoglobin at light wavelength 1 (Infrared)
BO2 is the extinction coefficient for oxygenated hemoglobin at light wavelength 2 (red) PA1 BR1 is the extinction coefficient for reduced hemoglobin at light wavelength 1 PA1 BR2 is the extinction coefficient for reduced hemoglobin at light wavelength 2 PA1 light wavelength 1 is infrared light PA1 light wavelength 2 is red light PA1 and R is the ratio of the optical density of wavelength 2 to wavelength 1 and is calculated as: ##EQU2## wherein I.sub.max2 is the maximum light transmitted at light wavelength 2 PA1 I.sub.min2 is the minimum light transmitted at light wavelength 2 PA1 I.sub.max1 is the maximum light transmitted at light wavelength 1 PA1 I.sub.min1 is the minimum light transmitted at light wavelength 1
The various extinction coefficients are determinable by empirical study as are well known to those of skill in the art. For convenience of calculation, the natural log of the ratios may be calculated by use of the Taylor expansion series for the natural log.
Several alternate methods of processing and interpreting optical signal data have been disclosed in the patents and references cited above.
Normally, the relative oxygen content of the patient's arterial pulses remains about the same from pulse to pulse and the average background absorption between pulses remains about the same. Consequently, the red and infrared light that is transmitted through the pulsatile flow produces a regularly modulated plethysmograph waveform having periodic optical pulses of comparable shape and amplitude and a steady state background transmittance. This regular pulse provides for an accurate determination of the oxygen saturation of the blood based on the detected relative maximum and minimum transmittance of the red and infrared light.
Changes in the patient's local blood volume at the optical detection site affect the absorption of light. These localized changes often result from motion artifact or respiratory artifact which introduce artificial pulses into the blood flow. For example, on each inhalation, the venus return is occluded slightly, which results in the background intensity component of transmittance being decreased due to the relatively larger volume of blood at the optical detection site. Exhalation allows the venus return to expand, thereby decreasing the volume of blood and increasing the background intensity component of transmittance. Consequently, the periodic optical pulses ride on a background intensity component of transmittance that rises and falls with blood volume change. This background intensity component variation, which is not necessarily related to changes in saturation, affects the pulse to pulse uniformity of shape, amplitude and expected ratio of the maximum to minimum transmittance, and can affect the reliability and accuracy of the saturation determination.
In addition, there are times when the patient's background level of oxygen saturation undergoes transient changes, for example, when the patient loses or reacquires oxygen exchange in the lungs while under gaseous anesthesia. Consequently, the detected red and infrared light transmittance changes and the detected plethysmograph waveform rises or falls over time with changes in the average oxygen saturation level in the patient's blood. The transient waveform distorts the pulse shape, amplitude, and the expected ratio of the pulses, which in turn affects the reliability and accuracy of the saturation determination.
Heretofore, with the foregoing known techniques for calculating arterial oxygen saturation, it was known that, during changes in the background intensity absorption component due to artifacts from changes in the patient's blood volume or transient saturation changes, the determined saturation value was not accurate and that it would not become accurate again until the average absorption (or transmittance) level stabilized at the end of the artifact or the saturation transient.
It also was known that saturation calculations based upon transient optical signals provided an over-estimation or under-estimation of the actual saturation value, depending upon the trend. The transmittance of red light near the 660 nanometer wavelength increases as oxygen saturation increases. This results in the detected optical signal value having a smaller pulsatile amplitude, i.e., a smaller relative difference between the maximum and minimum of the pulse. In contrast, the transmittance of the infrared light near the 910 nanometer wavelength decreases as saturation increases, which causes the infrared pulsatile amplitude--relative maximum to minimum--to increase. For both wavelengths, the transmittance changes with changing saturated are substantially linear and continuous in the range of clinical interest, i.e., oxygen saturations between 50% and 100%.
The accuracy of the estimation is of particular concern during rapid desaturation, where average oxygen saturation drops rapidly, but the saturation determination based on the detected optical signals indicates a greater drop than has actually occurred. The determined saturation thus may actuate low limit saturation alarms on an oximeter device that can result in unnecessary and wasteful efforts to resuscitate a patient not in danger.
Applicants believe that the change in transmittance that occurs between the maximum transmittance time and minimum transmittance time is due to the difference in arterial pulsatile length of pulse that has the same oxygen saturation. Because the pulsatile amplitude is quite small, typically less than 5% of the overall intensity change, any small change in overall or background transmittance, such as slight changes in average blood saturation, can have a relatively large effect in the difference in maximum and minimum intensity of the light levels. Because the transmittance effect of changing oxygen saturation is opposite in direction for the red light at 660 nanometers than for infrared light at 910 nanometers, this can result in over-estimation of the pulsatile ratio during periods when saturation is decreasing, and under-estimation during periods when saturation is increasing.
It is therefore an object of this invention to provide a method and apparatus for compensating for the effects of transient conditions in the actual optically detected signal, thereby providing a more accurate estimation of the actual oxygen saturation value.
It is another object of this invention to compensate for the effects of distortion in the detected oxygen saturation signal caused by artifacts due to localized blood volume changes.
It is another object of this invention to compensate for the effects of distortion in the detected oxygen saturation signal caused by transient saturation or blood volume artifact by using a determined rate of change from pulse to pulse, including using interpolation techniques.
It is another object of this invention to compensate for the effects of distortion in the detected oxygen saturation signal caused by transient saturation or blood volume artifact by using the low frequency characteristics of the detected signal values.